Microencapsulation technology holds promise in many areas of medicine. For example, some important applications are treatment of diabetes (Goosen, et al., 1985), production of biologically important chemicals (Omata, et al., 1979), evaluation of anti-human immunodeficiency virus drugs (McMahon, et al., 1990), encapsulation of hemoglobin for red blood cell substitutes, and controlled release of drugs. During encapsulation using prior methods, cells are often exposed to processing conditions which are potentially cytotoxic. These conditions include heat, organic solvents and non-physiological pH which can kill or functionally impair cells. Proteins are often exposed to conditions which are potentially denaturing and can result in loss of biological activity.
Further, even if cells survive processing conditions, the stringent requirements of encapsulating polymers for biocompatibility, chemical stability, immunoprotection and resistance to cellular overgrowth, restrict the applicability of prior art methods. For example, the encapsulating method based on ionic crosslinking of alginate (a polyanion) with polylysine or polyornithine (polycation) (Goosen, et al., 1987) offers relatively mild encapsulating conditions, but the long-term mechanical and chemical stability of such ionically crosslinked polymers remains doubtful. Moreover, these polymers when implanted in vivo, are susceptible to cellular overgrowth (McMahon, et al., 1990) which restricts the permeability of the microcapsule to nutrients, metabolites, and transport proteins from the surroundings. This has been seen to possibly lead to starvation and death of encapsulated islets of Langerhans cells (O'Shea and Sun, 1986).
Thus, there is a need for a relatively mild cell encapsulation method which offers control over properties of the encapsulating polymer. The membranes must be non-toxically produced in the presence of cells, with the qualities of being permselective, chemically stable, and very highly biocompatible. A similar need exists for the encapsulation of biological materials other than cells and tissues.
Biocompatibility
Synthetic or natural materials intended to come in contact with biological fluids or tissues are broadly classified as biomaterials. These biomaterials are considered biocompatible if they produce a minimal or no adverse response in the body. For many uses of biomaterials, it is desirable that the interaction between the physiological environment and the material be minimized. For these uses, the material is considered "biocompatible" if there is minimal cellular growth on its surface subsequent to implantation, minimal inflammatory reaction, and no evidence of anaphylaxis during use. Thus, the material should elicit neither a specific humoral or cellular immune response nor a nonspecific foreign body response.
Materials which are successful in preventing all of the above responses are relatively rare; biocompatibility is more a matter of degree rather than an absolute state. The first event occurring at the interface of any implant with surrounding biological fluids is protein adsorption (Andrade, et al., 1986). In the case of materials of natural origin, it is conceivable that specific antibodies for that material exist in the repertoire of the immune defense mechanism of the host. In this case a strong immune response can result. Most synthetic materials, however, do not elicit such a reaction. They can either activate the complement cascade or adsorb serum proteins which mediate cell adhesion, called cell adhesion molecules (CAMs) (Buck, et al., 1987). The CAM family includes proteins such as fibronectin, vitronectin, laminin, von Willebrand factor, and thrombospondin.
Proteins can adsorb on almost any type of material. They have positively and/or negatively charged regions, as well as hydrophilic and hydrophobic regions. They can thus interact with implanted material through any of these various regions, resulting in cellular proliferation at the implant surface. Complement fragments such as C3b can be immobilized on the implant surface and act as chemoattractants. They in turn can activate inflammatory cells such as macrophages and neutrophils and cause their adherence and activation on the implant. These cells attempt to degrade and digest the foreign material.
In the event that the implant is nondegradable and is too large to be ingested by large single activated macrophages, the inflammatory cells may undergo frustrated phagocytosis. Several such cells can combine to form foreign body giant cells. In this process, these cells release peroxides, hydrolytic enzymes, and chemoattractant and anaphylactic agents such as interleukins, which increase the severity of the reaction. They also induce the proliferation of fibroblasts on foreign surfaces.
Fibroblasts secrete a collagenous matrix which ultimately results in encasement of the entire implant in a fibrous envelope. Cell adhesion can also be mediated on a charged surface by the cell surface proteoglycans such as heparin sulfate and chondroitin sulfate (van Wachem, et al., 1987). In such a process, intermediary CAMs are not required and the cell surface can interact directly with the surface of the implant.
Enhancing Biocompatibility
Past approaches to enhancing biocompatibility of materials started with attempts at minimization of interfacial energy between the material and its aqueous surroundings. Similar interfacial tensions of the solid and liquid were expected to minimize the driving force for protein adsorption and this was expected to lead to reduced cell adhesion and thrombogenicity of the surface. For example, Amudeshwari et al. used collagen gels cross-linked in the presence of HEMA and MMA (Amudeshwari, et al., 1986). Desai and Hubbell showed a poly(HEMA)-MMA copolymer to be somewhat non-thrombogenic (Desai, N. P. and Hubbell, 1989).
Protein adsorption and desorption, however, is a dynamic phenomenon, as seen in the Vroman effect. This effect is the gradual displacement of one serum protein by another, through a well-defined series, until only virtually irreversibly adsorbed proteins are present on the surface. Affinity of protein in a partially dehydrated state for the polymer surface has been proposed as a determining factor for protein adsorption onto a surface (Baier, 1990). Enhancement of surface hydrophilicity has resulted in mixed success; increased hydrophilicity or hydrophobicity does not have a clear relation with biocompatibility (Coleman, et al., 1982; Hattori, et al., 1985). In some cases, surfaces with intermediate hydrophilicities demonstrate proportionately less protein adsorption. The minimization of protein adsorption may depend both upon hydrophilicity and the absence of change, as described further below, perhaps in addition to other factors.
Use of Gels in Biomaterials
Gels made of polymers which swell in water such as poly (HEMA), water-insoluble polyacrylates, and agarose, have been shown to be capable of encapsulating islet cells and other animal tissue (Iwata, et al., 1989; Lamberti, et al., 1984). However, these gels have undesirable mechanical properties. Agarose forms a weak gel, and the polyacrylates must be precipitated from organic solvents, thus increasing the potential for cytotoxicity. Dupuy et al. (1988) have reported the microencapsulation of islets by polymerization of acrylamide to form polyacrylamide gels. However, the polymerization process, if allowed to proceed rapidly to completion, generates local heat and requires the presence of toxic cross-linkers. This usually results in mechanically weak gels whose immunoprotective ability has not been established. Moreover, the presence of a low molecular weight monomer is required which itself is cytotoxic.
Microcapsules formed by the coacervation of alginate and poly(L-lysine) have been shown to be immunoprotective e.g., O'Shea et al., 1986. However, implantation for periods up to a week has resulted in severe fibrous overgrowth on these microcapsules (McMahon, et al. 1990; O'Shea, et al., 1986).
Use of PEO in Biomaterials
The use of poly(ethylene oxide) (PEO) to increase biocompatibility is well documented in the literature. The presence of grafted PEO on the surface of bovine serum albumin has been shown by Abuchowski et al. (1977) to reduce immunogenicity in a rabbit and to increase circulation times of exogenous proteins in animals. The biocompatibility of algin-poly(L-lysine) microcapsules has been significantly enhanced by incorporating a graft copolymer of PLL and PEO on the microcapsule surface (Sawhney, et al., in press)
The grafting of methoxy PEO onto polyacrylonitrile surfaces was seen by Miyama et al. (1988) to render the polyacrylonitrile surface relatively non-thrombogenic. Nagoaka et al. synthesized a graft copolymer of methacrylates with PEO and found the resulting polymer to be highly non-thrombogenic. Desai and Hubbell have immobilized PEO on poly(ethylene terepthalate) surfaces by forming a physical interpenetrating network (Desai et al., 1992); they have shown these surface to be highly resistant to thrombosis (Desai et al, 1991) and to both mammalian and bacterial cell growth (Desai, et al., submitted).
PEO is a unique polymer in terms of structure. The PEO chain is highly water soluble and highly flexible. Polymethylene glycol, on the other hand, undergoes rapid hydrolysis, while polypropylene oxide is insoluble in water. PEO chains have an extremely high motility in water and are completely non-ionic in structure. The synthesis and characterization of PEO derivatives which can be used for attachment of PEO to various surfaces, proteins, drugs etc. has been reviewed (Harris, 1985). Other polymers are also water soluble and non-ionic, such as poly(N-vinyl pyrrolidinone) and poly(ethyl oxazoline). These have been used to reduce interaction of cells with tissues. N. P. Desai et al. (1991). Water soluble ionic polymers, such as hyaluronic acid, have also been used to reduce cell adhesion to surfaces and can similarly be used.
Immobilization of PEO on a charged surface, such as a coacervated membrane of alginate-PLL, results in shielding of surface charges by the non-ionic PEO (Sawhney et al., in press). The highly motile PEO chain sweeps out a free volume in its microenvironment. The free volume exclusion effect makes the approach of a macromolecule (viz., a protein) close to a surface which has grafted PEO chains sterically unfavorable (Miyama, et al., 1988; Nagoaka, et al.; Desai, et al.; Sun, et al., 1987). Thus protein adsorption is minimized and cell adhesion is reduced, resulting in surfaces showing increased biocompatibility.
Immobilization of PEO on a surface has been largely carried out by the synthesis of graft copolymers having PEO side chains (Sawhney, et al.; Miyama, et al., 1988; Nagoaka, et 81.). This process involves the custom synthesis of monomers and polymers for each application. The use of graft copolymers, however, still does not guarantee that the surface "seen" by a macromolecule consists entirely of PEO.
Electron beam cross-linking has been used to synthesize PEO hydrogels, and these biomaterials have been reported to be non-thrombogenic (Sun, et al., 1987; Dennison, H. A., 1986). However, use of an electron beam precludes the presence of any living tissue due to the sterilizing effect of this radiation. Also, the networks produced are difficult to characterize due to he non-specific cross-linking induced by the electron beam.
Photopolymerizable PEG diacrylates have been used to entrap yeast cells for fermentation and chemical conversion (Kimura et al. 1981; Omata et al., 1981; Okada et al. 1987). However, yeast cells are widely known to be much hardier, resistant to adverse environments and elevated temperatures, and more difficult to kill when compared to mammalian cells and human tissues. For example, yeast may be grown anaerobically, whereas mammalian cells may not; yeast are more resistant to organic solvents (e.g., ethanol to 12%) than are mammalian cells (e.g., ethanol to &lt;1%); and yeast possess a polysaccharide cell wall, whereas mammalian cells, proteins, polysaccharides, and drugs do not. None of these references, however, discuss the exposure of sensitive eukaryotic tissue, organisms, or sensitive molecules to the chemical conditions used during polymerization because their polymerization conditions are incompatible with sensitive materials. For example, there are no reports of the encapsulation of mammalian cells using prior art photosensitive prepolymers without a marked loss of cellular function.
Other earlier encapsulations of cells within photopolymerizable materials have focused on microbial cells (Kimura et al., 1981; Omata et al., 1981; Okada et al., 1987; Tanaka et al., 1977; Omata et al 1979a; Omata et al., 1979b; Chun et al. 1981; Fukui et al., 1976; Fukui et al., 1984). Each of these reports, however, describes the use of near ultraviolet light (wavelength&lt;320 nm), which is injurious to more sensitive cells such as mammalian cells or higher eukaryotic cells. In the original presentation of the technique (Fukui et al., 1976), the authors state in the final sentence that the technique would be appropriate for microbial cells, but provide no indication of usefulness for more sensitive cells. In a more recent and complete review of the technique (Fukui et al., 1984), the authors, in section 6 entitled "Entrapped Living Cells" provide no teaching regarding cells other than microbial cells, and in section 7 entitled "Future Prospects" they also provide no such teaching.
Moreover, the prior use of such materials for the entrapment of biological materials is entirely focused on industrial technology, rather than biomedical technology. For example, no attention is paid to biocompatibility, including formulation of the gel to avoid the problems described above. This is an important issue, since bioincompatibility in biomedical applications leads to xenograft failure in therapeutically transplanted cells for the evaluation of drug efficacy (O'Shea et al., 1986) and to xenograft failure in diagnostically transplanted cells (McMahon et al., 1990). Similarly, bioincompatibility would lead to the failure of encapsulated enzymes (for example, therapeutic enzymes encapsulated and circulating or implanted in a blood-rich tissue). Such encapsulated and entrapped enzymes could leave the circulation by interaction with the reticuloendothelial system (Hunt et al., 1985) or could become overgrown with tissues in a foreign body reaction.
Other ways of producing PEO hydrogels include use of PEO chains end capped with n-alkane chains, which associate in aqueous media to form stable gels (Knowles, et al., 1990). No biological properties of these materials have been reported, however. Thus, the prior art contains no description of methods to form biocompatible PEO networks on three-dimensional living tissue surfaces without damaging encapsulated tissue.
Among the techniques for encapsulating mammalian tissue with polymers other than PEO is a method of photopolymerizing the monomer 2-hydroxyethyl methacrylate ("HEMA") and the crosslinking agent ethylene glycol dimethacrylate ("EGDA") in a cylindrical mold containing the biological material (Ronel, et al., 1981). The product of this reaction, a cylindrical gel with cells embedded throughout, is frozen and then finely ground into small particles. This technique, however, suffers from a number of disadvantages. First, because the cylindrical gel is broken along random planes, shearing will often occur through pockets of cells, leaving some cells exposed to the host immune system. Second, HEMA and EGDA are small cytotoxic molecules capable of penetrating the cellular membrane. Third, the resulting polymer membrane has uneven pore sizes which vary to an upper limit of 20 microns, thereby allowing transit of immune response molecules. These drawbacks are reflected in data which show that tissue remains viable for only 2-3 days after this encapsulation process.